The invention relates generally to diagnostic imaging and, more particularly, to an apparatus for reducing photodiode thermal gain coefficient in a photodiode array.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom.
Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
A CT detector typically has stringent specifications on the channel-to-channel or pixel-to-pixel differential signal error, especially for the center part of the detector ring. For example, the tolerance for pixel-to-pixel differential signal error may be as low as 200 ppm. One of the typical contributions to the differential signal error of the detector is from the photodiode arrays due to the existence of a diode thermal coefficient of gain (“gain tempco”) and due to temperature variations within and between diode arrays. To minimize this contribution, CT detectors are designed with low temperature variation at the diode arrays. With the increase of the detector size to provide more and more coverage, this thermal design becomes more and more challenging. Other methods for minimizing contributions due to temperature variations include thermal management/cooling systems designed to remove excess temperature from the CT detectors. These thermal management systems, however, are often bulky and add excess weight and complexity to the CT gantry system.
Therefore, it would be desirable to design an apparatus capable of low differential signal errors due to temperature variations within a photodiode array of any size.